Images of the interiors of bodies may be acquired using various types of tomographic techniques, which involve recording and measuring radiation from tissues and processing acquired data into images.
One of these tomographic techniques is positron emission tomography (PET), which involves determining spatial distribution of a selected substance throughout the body and facilitates detection of changes in the concentration of that substance over time, thus allowing to determine the metabolic rates in tissue cells.
The selected substance is a radiopharmaceutical administered to the examined object (e.g. a patient) before the PET scan. The radiopharmaceutical, also referred to as an isotopic tracer, is a chemical substance having at least one atom replaced by a radioactive isotope, e.g. 11C, 15O, 13N, 18F, selected so that it undergoes radioactive decay including the emission of a positron (antielectron). The positron is emitted from the atom nucleus and penetrates into the object's tissue, where it is annihilated in reaction with an electron present within the object's body.
The phenomenon of positron and electron annihilation, constituting the principle of PET imaging, consists in converting the masses of both particles into energy emitted as annihilation photons, each having the energy of 511 keV. A single annihilation event usually leads to formation of two photons that diverge in opposite directions at the angle of 180° in accordance with the law of conservation of the momentum within the electron-positron pair's rest frame, with the straight line of photon emission being referred to as the line of response (LOR). The stream of photons generated in the above process is referred to as gamma radiation and each photon is referred to as gamma quantum to highlight the nuclear origin of this radiation. The gamma quanta are capable of penetrating matter, including tissues of living organisms, facilitating their detection at certain distance from object's body. The process of annihilation of the positron-electron pair usually occurs at a distance of several millimeters from the place of the radioactive decay of the isotopic tracer. This distance constitutes a natural limitation of the spatial resolution of PET images to a few millimeters.
A PET scanner comprises detection devices used to detect gamma radiation as well as electronic hardware and software allowing to determine the position of the positron-electron pair annihilation event on the basis of the position and time of detection of a particular pair of the gamma quanta. The radiation detectors are usually arranged in layers forming a ring around object's body and are mainly made of an inorganic scintillation material. A gamma quantum enters the scintillator, which absorbs its energy to re-emit it in the form of light (a stream of photons). The mechanism of gamma quantum energy absorption within the scintillator may be of dual nature, occurring either by means of the Compton's effect or by means of the photoelectric phenomenon, with only the photoelectric phenomenon being taken into account in calculations carried out by current PET scanners. Thus, it is assumed that the number of photons generated in the scintillator material is proportional to the energy of gamma quanta deposited within the scintillator.
When two annihilation gamma quanta are detected by a pair of detectors at a time interval not larger than several nanoseconds, i.e. in coincidence, the position of annihilation point along the line of response may be determined, i.e. along the line connecting the detector centers or the points within the scintillator strips where the energy of the gamma quanta was deposited. The coordinates of annihilation place are obtained from the difference in times of arrival of two gamma quanta to the detectors located at both ends of the LOR. In the prior art literature, this technique is referred to as the time of flight (TOF) technique and the PET scanners utilizing time measurements are referred to as TOF-PET scanners. This technique requires that the scintillator has time resolution of a few hundred picoseconds.
Light pulses reaching the scintillator can be converted into electric pulses by means of photomultipliers or photodiodes. Electric signals from the converters carry information on positions and times of the annihilation quanta subject to detection, as well as on the energy deposited by these quanta.
The principal elements of the signal processing system within the radiation detectors are leading edge discriminators and constant fraction discriminators. These elements, combined with time-to-digital converters, facilitate the measurement of time at which the electric signals generated at these detectors exceed a preset reference voltage or a preset signal amplitude fraction, respectively. Said discriminators are built on the basis of standard electronic components and include, among other components, a current source, a preamplifier, a comparator, a shaper, capacitors, resistors, diodes, transistors and transmission lines. If the detector signal is higher than the threshold voltage set at the discriminator, a logical signal is generated at the discriminator output, carrying information on the time at which the gamma quantum was recorded. The charge is measured by means of analog-to-digital converters.
Temporal resolutions of leading edge and constant fraction discriminators are limited by the dependence of the discriminator response on the shape of signals and, in case of leading edge discriminators, also on the amplitude of input signals. Due to the so-called time walk effect, time determined using leading edge discriminators changes along with the signal amplitude. The effect may be adjusted to a certain degree if the signal charge or amplitude is measured simultaneously. In case of constant fraction discriminators, the time at which the signal exceeds the preset amplitude fraction is generally not dependent on the amplitude, but it may change depending on the shape of the signal (i.e on the temporal distribution of photons).
Logical signals generated at discriminators are processed by means of sequences of logical operations within a triggering system. These operations result in a logical signal providing information on whether the recorded event should be subjected to further electronic processing. The sequences of logical operations are selected depending on the types of detectors, configuration of modules and the frequencies of recorded events; the main objective of these operations is to discard signals that are not useful for image reconstruction and thus to minimize acquisition dead times as well as times required to process the data and reconstruct the images.
The PCT applications WO2011/008119 and WO2011/008118 describe various aspects of PET scanners that may be of relevance for understanding this description, in particular, a method for determining the place of ionization on the basis of the distribution of times or amplitudes of signals measured at different positions along the scintillator. These documents describe solutions that are based on the measurements of the times of flight required for light pulses to reach detector edges. Changes in shapes and amplitudes of signals depending on the place of ionization and the quantity of energy constitute a constraint in temporal resolutions that can be achieved using the technique. The larger the scintillator, the larger the variations in signal shapes and amplitudes. For the above reasons, temporal resolutions of less than 100 ps cannot be obtained for large scintillator blocks according to the prior art. Temporal resolution also impacts the resolution of ionization place determination. In case of polymer scintillators (preferred due to their low price), amplitudes of signals generated by the gamma quanta, including annihilation gamma quanta used in positron emission tomography, are characterized by continuous distribution resulting from interactions between gamma quanta and electrons occurring mostly via the Compton effect with a negligibly low probability of a photoelectric effect. As a consequence, signal amplitudes in polymer scintillators may change even if the signals originated in the same position. In case of Compton interactions, constraints in the achieved resolution are due to the fact that the amplitude of electric signals generated by the photomultipliers depends on two unknown values, namely on distance between the ionization place and the photomultiplier and on energy deposited by the gamma quantum. The effects described above contribute to deterioration in both temporal and spatial resolution also in case of monoenergetic energy-loss distributions, which occur e.g. in the photoelectric effect.
As evidenced by the shortcomings of the state of the art signal analysis techniques described above, there is a need to significantly improve temporal and spatial resolution of detectors being used in medical diagnostic techniques that require the recording of ionizing radiation. The need to improve resolution is particularly high in large-sized detectors.
One of the developed improvement methods is continuous sampling of analog signals in temporal domain. Continuous sampling known from the state of the art is associated with operation of an ADC converter which collects a specific number of analog signal samples at predefined time intervals. However, the method is not capable of improving results in case of rapid signals from polymer scintillators, characterized by decay and rise times on the order of 1 ns. Sampling frequencies that may be practically applied in devices featuring a large number of detectors are on the order of 100 MHz (Flash ADC). This sampling frequency corresponds to sampling intervals of 10 ns, which are comparable to the duration of the signal itself. Therefore, even if the sampling frequencies were higher by an order of magnitude, they would still be insufficient for analyzing signals from polymer scintillators.
It would be desirable to develop a detector and a method for determining the position and time of ionization in polymer scintillators.